U.S. patent number 8,197,547 [Application Number 11/179,387] was granted by the patent office on 2012-06-12 for radiovisible hydrogel intervertebral disc nucleus.
This patent grant is currently assigned to Howmedica Osteonics Corp.. Invention is credited to Christopher DeMaria, Paul Higham, Chau Ngo, Philip F. Williams, III.
United States Patent |
8,197,547 |
Higham , et al. |
June 12, 2012 |
Radiovisible hydrogel intervertebral disc nucleus
Abstract
A spinal implant for replacing the natural nucleus of the disc
made from a polymer such as hydrogel having a radiopaque material
located within the polymer. The material may be in the form of a
powder dispersed throughout the polymer or may be in he form of a
powder dispersed in layers or in other specific locations within
the polymer. The radiopaque material is metal such as gold,
tungsten, titanium, tantalum or platinum. The metal may also be in
the form of a foil or wire located within the hydrogel such as
polyurethane, thereby making the implant visible on x-rays. Other
polymers besides hydrogel may be used with the radiopaque material
being dispersed therein.
Inventors: |
Higham; Paul (Ringwood, NJ),
Ngo; Chau (Secaucus, NJ), DeMaria; Christopher (Glen
Rock, NJ), Williams, III; Philip F. (Teaneck, NJ) |
Assignee: |
Howmedica Osteonics Corp.
(Mahwah, NJ)
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Family
ID: |
31991880 |
Appl.
No.: |
11/179,387 |
Filed: |
July 12, 2005 |
Prior Publication Data
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Document
Identifier |
Publication Date |
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US 20050267583 A1 |
Dec 1, 2005 |
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Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
Issue Date |
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10244306 |
Sep 16, 2002 |
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Current U.S.
Class: |
623/17.16 |
Current CPC
Class: |
A61L
27/50 (20130101); A61F 2/442 (20130101); A61L
27/52 (20130101); A61F 2002/30971 (20130101); A61F
2310/00131 (20130101); A61F 2/441 (20130101); A61F
2310/00155 (20130101); A61F 2/30965 (20130101); A61L
2430/38 (20130101); A61F 2002/30616 (20130101); A61F
2310/00137 (20130101); A61F 2002/3008 (20130101); A61F
2250/0098 (20130101); A61F 2310/00149 (20130101); A61F
2310/00023 (20130101); A61F 2002/444 (20130101); A61F
2250/0064 (20130101) |
Current International
Class: |
A61F
2/44 (20060101) |
Field of
Search: |
;623/17.11,17.12,17.16
;606/77 ;523/117 |
References Cited
[Referenced By]
U.S. Patent Documents
Foreign Patent Documents
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30 18 966 |
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Dec 1981 |
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DE |
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0 462 512 |
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Dec 1991 |
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EP |
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9726847 |
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Jul 1997 |
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WO |
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WO-98/55053 |
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Dec 1998 |
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WO |
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WO-99/62439 |
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Dec 1999 |
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WO |
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WO-00/03691 |
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Jan 2000 |
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WO |
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WO-02/17825 |
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Mar 2002 |
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WO |
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WO-03/045274 |
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Jun 2003 |
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WO |
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Other References
North American Spine Society Proceedings, 16th Annual Meeting, Oct.
31, 2001, p. 178-179. cited by other .
Horak et al. "New Radiopaque poly-HEMA based hydrogel particles"
1997, vol. 34, pp. 183-188. cited by other .
Thanoo et al. "Radiopaque hydrogel micrsperes" 1988, vol. 6. No. 2,
p. 233-244. cited by other.
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Primary Examiner: Pellegrino; Brian E.
Attorney, Agent or Firm: Lerner, David, Littenberg, Krumholz
& Mentlik, LLP
Claims
The invention claimed is:
1. A spinal implant for replacing the natural nucleus of a disc
comprising: a layered hydrogel material with at least one layer of
the hydrogel material having powdered radiopaque material therein,
the implant having an uppermost and lowermost surface and outer
side surfaces which contact a disc annulus extending therebetween;
and the powdered radiopaque material incorporated into the at least
one layer of the hydrogel material in the form of at least one
discrete planar layer, an upper and lower side of the at least one
discrete planar layer including radiopaque powder contacts a layer
of the hydrogel material not containing radiopaque powder, the at
least one planar layer containing radiopaque powder extends the
entire length of the layer to the outer side surfaces of the
implant which contact the disc annulus, the powdered radiopaque
material in the at least one hydrogel layer having radiopaque
material are in a concentration of less than 0.5 gram of radiopaque
powdered material to 1 gram of hydrogel and moves with the hydrogel
body as the body changes dimensions and geometry under load, the
upper and lowermost layers of the implant having no radiopaque
powder.
2. The spinal implant as set forth in claim 1 wherein there are
between 2 and 5 discrete planar layers of the hydrogel containing
the radiopaque powder in the implant body oriented in parallel
planes between the upper and lowermost layers of the implant.
3. The spinal implant as set forth in claim 2 wherein the
radiopaque material is metal particles.
4. The spinal implant as set forth in claim 3, wherein the metal
particles are selected from the group consisting of gold, tungsten,
tantalum, platinum, titanium and a combination thereof.
5. The spinal implant as set forth in claim 4, wherein the
particles have a diameter of between 10 and 100 .mu.m.
6. The spinal implant as set forth in claim 5, wherein the
particles have a maximum diameter of about 75 .mu.m.
7. The spinal implant as set forth in claim 1 wherein the
radiopaque material is metal particles.
8. The spinal implant as set forth in claim 7, wherein the metal
particles are selected from the group consisting of gold, tungsten,
tantalum, platinum, titanium and a combination thereof.
9. The spinal implant as set forth in claim 8, wherein the
particles have a diameter of between 10 and 100 .mu.m.
10. The implant as set forth in claim 9, wherein the metal
particles have a maximum diameter of about 75 .mu.m.
11. A polymeric implant for replacement of a part of a natural disc
nucleus comprising a layered hydrogel body comprising a hydrogel
material, radiopaque particles are dispersed in at least two
discrete planar layers of the hydrogel body, the hydrogel body
having uppermost and lowermost surfaces and an outer side surface
connecting the upper and lower surfaces, the side surface for
contacting a disc annulus, the two discrete planar layers of the
hydrogel body having radiopaque particles extending the entire
length of the layer to the outer side surface of the layered
hydrogel body contacting the disc annulus surrounding the nucleus,
each discrete layer containing radiopaque particles having upper
and lower surfaces adjacent a layer of the hydrogel material
without radiopaque particles, the upper and lowermost surfaces of
the implant having no radiopaque powder the radiopaque particles in
the at least two hydrogel material layers in a concentration less
than 0.5 gram of radiopaque particles to 1 cubic centimeter of the
hydrogel and the particles move with the layered hydrogel body
structure as the implant changes geometry under load.
12. The spinal implant as set forth in claim 11 wherein there are
between 3 and 5 layers of the hydrogel material containing
radiopaque particles.
13. The spinal implant as set forth in claim 12 wherein the
radiopaque particles are a metal.
14. The spinal implant as set forth in claim 13, wherein the metal
particles are selected from the group consisting of gold, tungsten,
tantalum, platinum, titanium and a combination thereof.
15. The spinal implant as set forth in claim 14, wherein the
particles have a diameter of between 10 and 100 .mu.m.
16. The spinal implant as set forth in claim 15, wherein the
particles have a maximum diameter of about 75 .mu.m.
17. The spinal implant as set forth in claim 11 wherein the layers
are planar.
18. A spinal implant for replacement of a part of a natural disc
nucleus comprising: a hydrogel structure made from layers of gelled
polymeric solution, part of the hydrogel structure having powdered
radiopaque material within the gelled polymeric solution wherein
the implant has a plurality of discrete planar layers of the gelled
polymeric solution including the powdered radiopaque material and a
plurality of planar layers of the gelled polymeric solution without
the powdered radiopaque material wherein at least one discrete
planar layer including radiopaque material contacts a surface of
the planar layer of the gelled polymeric solution not containing
radiopaque material on first and second planar sides of the layer
containing the radiopaque material, the at least one planar layer
including radiopaque material extending the entire length of the
layer between outer side surfaces of the implant which contact a
disc annulus, an uppermost and lowermost surface of the implant
made of a layer of gelled polymeric solution without the radiopaque
material, wherein the radiopaque material is in the form of
radiopaque particles which particles are in the gelled polymeric
solution layer in a concentration less than 0.5 gram of radiopaque
particles to 1 cubic centimeter of the polymeric solution and the
particles move with the hydrogel structure as the implant changes
geometry under load.
19. The spinal implant as set forth in claim 18 wherein the
radiopaque material particles are metal particles of between 10
.mu.m and 100 .mu.m in size.
20. The spinal implant as set forth in claim 19 wherein the metal
particles are selected from the group consisting of gold, tungsten,
tantalum, platinum, titanium and a combination thereof.
21. The spinal implant as set forth in claim 18 wherein there are
three planar layers of gelled polymeric solution containing
radiopaque material.
Description
BACKGROUND OF THE INVENTION
This application is a divisional of U.S. application Ser. No.
10/244,306, filed on Sep. 16, 2002, the disclosure of which is
incorporated herein by reference.
This invention relates to a prosthetic intervertebral disc nucleus.
More particularly, it relates to an artificial disc nucleus made of
a hydrogel material having a radiovisible material therein.
The intervertebral disc is a complex joint anatomically and
functionally. It is composed of three component structures; the
nucleus pulposus (the nucleus), the annulus fibrosus (the annulus)
and the vertebral end-plates. The biochemical composition and
anatomical arrangements within these component structures are
related to the biomechanical function of the disc.
The nucleus occupies about 25-40% of the total disc cross-sectional
area. It is primarily composed of mucoid material containing mainly
proteoglycans with a small amount of collagen. The proteoglycans
consist of a protein core with chains of negatively charges keratin
sulphate and chondroitin sulphate covalently attached thereto. Due
to these constituents, the nucleus is a loose hydrogel which
usually contains about 70-90% water by weight. Although the nucleus
plays an important role in the biomechanical function of the disc,
the mechanical properties of the disc are not well known, largely
because of the loose hydrogel nature of the nucleus.
As the nucleus is surrounded by the annulus and vertebral
end-plates and the negatively charged sulphate groups are
immobilized due to the attachment of these groups to the polymer
matrix, the matrix has a higher concentration of counter ions than
its surroundings. This ion concentration results in a higher
osmotic pressure than the annulus e.g., ranging from about 0.1 to
about 0.3 MPa. As a result of the high fixed charge density of the
proteoglycan the matrix exerts an osmotic swelling pressure which
can support an applied load in much the same way as air pressure in
a tire supports the weight of a car.
It is the osmotic swelling pressure and hydrophilicity of the
nucleus matrix that offers the nucleus the capability of imbibing
fluid until it is balanced with the internal resistance stresses,
due to the tensile forces of the collagen network, and the external
stresses due to the loads that are applied by muscle and ligament
tension. The swelling pressure (Ps) of the nucleus is directly
dependent on the concentration and fixed charge densities of
proteoglycan, i.e., the higher the concentration and fixed charge
densities of proteoglycan the higher will be the swelling pressure
of the nucleus. The external pressure changes with posture. When
the human body is supine the compressive load on the third lumbar
disc is 300 newton (N) which rises to 700 N when an upright stance
is assumed. The compressive load increases, yet again, to 1200 N
when the body is bent forward by only 20.degree. C. When the
external pressure (Pa) increases the previous balance, i.e. Ps=Pa,
is upset. To reach a new balance, the swelling pressure has to
increase. This increase is achieved by increasing the proteoglycan
concentration in the nucleus which is achieved by reducing the
fluid in the nucleus. That is why discs lose about 10% of their
height, as a result of creep, during the daytime. When the external
load is released i.e., Ps is greater than Pa, the nucleus will
imbibe fluid from its surroundings in order to reach the new
equilibrium value. It is this property of the nucleus that is
mainly responsible for the compressive properties of the disc.
The annulus forms the outer limiting boundary of the disc. It is
composed of highly structured collagen fibers embedded in an
amorphous base substance which is also composed of water and
proteoglycans. The amount of proteoglycans is lower in the annulus
than in the nucleus. The collagen fibers of the annulus are
arranged in concentric laminated bands or lamella, (about 8-12
layers thick) with a thicker anterior wall and thinner posterior
wall. In each lamella, the fibers are parallel and attached to the
superior and inferior vertebral bodies at an angle of about
30.degree. form the horizontal plane of the disc in both
directions. This design particularly resists twisting because the
half of the fibers cocked in one direction will tighten as the
vertebrae rotate relative to each other in the other direction. The
composition of the annulus along the radial axis is not uniform.
There is a steady increase in the proportion of collagen from the
inner to the outer sections of the annulus. This difference in
composition may reflect the need of the inner and outer regions of
the annulus to blend into very different tissues while maintaining
the strength of the structure. Only the inner lamellae are anchored
to the end-plates forming an enclosed vessel for the nucleus. The
collagen network of the annulus restrains the tendency of the
nucleus gel to absorb water from surrounding tissues and swell.
Thus, the collagen fibers in the annulus are always in tension, and
the nucleus gel is always in compression.
The two vertebral end-plates are composed of hyaline cartilage,
which is a clear, "glassy" tissue, that separates the disc from the
adjacent vertebral bodies. This layer acts as a transitional zone
between the hard, bony vertebral bodies and the soft disc. Because
the intervertebral disc is avascular, most nutrients that the disc
needs for metabolism are transported to the disc by diffusion
through the end-plate area.
The intervertebral joint exhibits both elastic and viscous
behavior. Hence, during the application of a load to the disc there
will be an immediate "distortion" or "deformation" of the disc,
often referred to as "instantaneous deformation." It has been
reported that the major pathway by which water is lost, from the
disc during compression, is through the cartilage end-plates. Since
the water permeability of the end-plates is in the range of about
0.20 to about 0.85.times.10.sup.-17m.sup.4N.sup.-1 sec.sup.-1 it is
reasonable to assume that under loading, the initial volume of the
disc is constant while the load is applied. Because the natural
nucleus of the disc is in the form of a loose hydrogel, i.e., a
hydrophilic polymeric material which is insoluble in water, it can
be deformed easily, the extent of deformation of the disc being
largely dependent on the extensibility of the annulus. It is
generally believed that hydrostatic behavior of the nucleus plays
an important role in the normal static and dynamic load-sharing
capability of the disc and the restoring force of the stretched
fibers of the annulus balances the effects of the nucleus swelling
pressure. Without the constraint by the annulus, annular bulging of
the nucleus would increase considerably. If the load is maintained
at a constant level, a gradual change in joint height, commonly
referred to as "creep" will occur as a function of time.
Eventually, the creep will stabilized and the joint is said to be
in "equilibrium." When the load is removed the joint will gradually
"recover" to its original height before loading. The creep and
relation rates depend on the amount of load applied, the
permeability of the end-plates and the water binding capability of
the nucleus hydrogel. Creep and relaxation are essential processes
in pumping fluid in and out of the disc.
Degeneration of the intervertebral disc is believed to be a common
cause of final pathological changes and back pain. As the
intervertebral disc ages it undergoes degeneration. The changes
that occur are such that, in many respects, the composition of the
nucleus seems to approach that of the inner annulus. Intervertebral
disc degeneration is, at least in part, the consequence of
compositional changes in the nucleus. It has been found that both
the molecular weight and the amount of proteoglycans in the nucleus
decrease with age, especially in degenerated discs, and the ratio
of keratin sulphate to chondroitin sulphate in the nucleus
increases. This increase in the ratio of keratin sulphate to
chondroitin sulphate and decrease in proteoglycan content decreases
the fixed charge density of the nucleus from about 0.28 meq/ml to
about 0.18-0.20 meq/ml. These changes cause the nucleus to lose
part of its water binding capability which decreases the maximum
swelling pressure it can exert. As a result, the maximum water
content drops from over about 85%, in preadolescence, to about
70-75% in middle age. The glycosaminoglycan content of prolapsed
discs has been found to be lower, and the collagen content higher,
than that of normal discs of a comparable age. Discs L-4-L-5 and
L-5-S-1 are usually the most degenerated discs.
It is known that although the nucleus only occupies about one third
of the total disc area, it takes about 70% of the total loading in
a normal disc. Thus, it has been found that the compressive load on
the nuclei of moderately degenerated discs is about 30% lower than
in comparable normal discs but the compressive load on the annulus
increases by 100% in the degenerated discs. This load change is
primarily caused by the structural changes in the disc as discussed
above. The excess load on the annulus, of the degenerated disc,
causes reduction of the disc height and excessive movement of the
spinal segments. The flexibility of the disc produces excessive
movement of the collagenous fibers which in turn, injures the fiber
attachments and causes delamination of the well organized fibers of
the annulus ring. The delamination annulus can be further weakened
by stress on the annulus and in severe cases this stress will cause
tearing of the annulus. This whole process is very similar to
driving on a flat tire, where the reinforcement layer will
eventually delaminate. Because the thickness of the annulus is not
uniform, with the posterior portions being thinner than the
anterior portions, delamination and lesions usually occur in the
posterior area first.
The spinal disc may also be displaced or damaged due to trauma or
diseases. In these cases, and in the case of disc degeneration, the
nucleus may herniate and/or protrude into the vertebral canal or
intervertebral foramen, in which case it is known as a herniated or
"slipped" disc. This disc may in turn press upon the spinal nerve
that exits the vertebral canal through the partially obstructed
foramen, causing pain or paralysis in the area of its distribution.
The most frequent site of occurrence of a herniated disc is in the
lower lumbar region. A disc herniation in this area often involves
the inferior extremities by compressing the sciatic nerve.
There are basically three types of treatment currently being used
for treating low back pain caused by injured or degenerated discs:
conservative care, discectomy and fusion. Each of these treatments
has its advantages and limitations. The vast majority of patients
with low back pain, especially those with first time episodes of
low back pain, will get better with conservative treatment.
However, it is not necessarily true that conservative care is the
most efficient and economical way to solve the low back pain
problem.
Discectomy usually provides excellent short term results in
relieving the clinical symptoms, by removing the herniated disc
material, usually the nucleus, which causes the low back pain
either by compressing the spinal nerve or by chemical irritation.
Clearly, a discectomy is not desirable from a biomechanical point
of view. In a healthy disc, the nucleus takes the most
compressional load and in a degenerated disc this load is primarily
distributed onto the annulus ring which, as described above, causes
tearing and delamination of the annulus. Removal of the nucleus in
a discectomy actually causes distribution the compressive load onto
the annulus ring thereby narrowing the disc spaces. It has been
reported that a long-term disc height decrease might be expected to
cause irreversible osteoarthritis-like changed in the facet joint.
That is why discectomy yields poor long term benefits and results
in a high incidence of reherniation.
Fusion generally does a good job in eliminating symptoms and
stabilizing the joint. However, because the motion of the fused
segment is restricted, the range of motion of the adjoining
vertebral discs is increased possibly enhancing their degenerative
processes.
Because of these disadvantages, it is desirable to use a prosthetic
joint device which not only is able to replace the injured or
degenerated intervertebral disc, but also can mimic the
physiological and the biomechanical function of the replaced disc
and prevent further degeneration of the surrounding tissue.
Artificial discs are well known in the prior art. U.S. Pat. No.
3,867,728, to Stubstad et al., relates to a device which replaces
the entire disc. This device is made by laminating vertical,
horizontal or axial sheets of elastic polymer. U.S. Pat. No.
4,309,777, to Patil, relates to a prosthetic utilizing metal
springs and cups. A spinal implant comprising a rigid solid body
having a porous coating on part of its surface is shown in Kenna's
U.S. Pat. No. 4,714,469. An intervertebral disc prosthetic
consisting of a pair of rigid plugs to replace the degenerated disc
is referred by Kuntz, U.S. Pat. No. 4,349,921. U.S. Pat. Nos.
4,772,287 and 4,904,260 to Ray et al., teach the use of a pair of
cylindrical prosthetic intervertebral disc capsules with or without
therapeutical agents. U.S. Pat. No. 4,911,718 to Lee et al.,
relates to an elastomeric disc spacer comprising three different
parts; nucleus, annulus and end-plates, of different materials. At
the present time, none of these inventions has become a product in
the spinal care market. Bao et al., in U.S. Pat. Nos. 5,047,055 and
5,192,326 (assigned to the assignee of this invention and
incorporated herein by reference) describe artificial nuclei
comprising hydrogels in the form of large pieces shaped when fully
hydrated, to generally conform to the disc cavity or hydrogel beads
within a porous envelope, respectively. The hydrogels have an
equilibrium water content (EWC) of at least about 30% and a
compressive strength of at least about 1 meganewtons per square
meter (1 MNm.sup.-2) when subjected to the constraints of the
annulus and end plates of the disc. Preferably, the compressive
strength of the nucleus is about 4 MNm.sup.-2 or even higher.
The primary disadvantage of the invention of Substad et al., Patil,
Kenna and Lee et al., is that use of their prosthesis requires
complete replacement of the natural disc which involves numerous
surgical difficulties. Secondly, the intervertebral disc is a
complex joint, anatomically and functionally, comprising the
aforementioned three component structures, each of which has its
own unique structural characteristics. Designing and fabricating
such a complicated prosthesis from acceptable materials, which will
mimic the function of the natural disc, is very difficult. A
further problem is the difficulty of preventing the prosthesis from
dislodging. Fourthly, even for prostheses which are only intended
for replacing the nucleus, a major obstacle has been to find a
material which is similar to the natural and is also able to
restore the normal function of the nucleus. Hydrophobic elastomers
and thermoplastic polymers are not desirable for use in the
prosthetic nuclei due to their significant inherent differences
from the natural nucleus e.g., lack of hydrophilicty, in the
elastomers, and lack of flexibility in their thermoplasts.
These problems are not solved by Kuntz, who uses elastic rubber
plugs, or by Froning and Ray et al., who use bladders, or capsules,
respectively, which are filled with a fluid or thixotropic gel.
According to the Ray and Froning patents, liquid was used to fill
the capsules and bladders, respectively, thereby requiring that
their membranes be completely sealed to prevent fluid leakage. As a
consequence, those devices cannot completely restore the function
of the nucleus which allows body fluid to diffuse in and out during
cyclic loading thereby providing the nutrients the disc needs.
The Bao et al., prosthetic lumbar disc nuclei are made from
hydrogels. Hydrogels have been used in biomedical applications,
such as contact lenses. Among the advantages of hydrogels is that
they are more biocompatible than hydrophobic elastomers and metals.
This biocompatibility is largely due to the unique characteristics
of hydrogels in that they are soft and contain water like the
surround tissues and have relatively low frictional coefficients
with respect to the surrounding tissues. The biocompatibility of
hydrogels results in prosthetic nuclei which are more easily
tolerated in the body. Furthermore, hydrophobic elastomeric and
metallic gels will not permit diffusion of aqueous compositions,
and their solutes, therethrough.
An additional advantage of some hydrogels is their good mechanical
strength which permits them to withstand the load on the disc and
restore the normal space between the vertebral bodies. The
aforementioned nuclei of Bao et al. have high mechanical strength
and are able to withstand the body loads and assist in the healing
of the defective annuli.
Other advantages of the hydrogels, used in Bao et al. nuclei, are
their excellent viscoelastic properties and shape memory. Hydrogels
contain a large amount of water which acts as a plasticizer. Part
of the water is available as free water which has more freedom to
leave the hydrogel when the hydrogel is partially dehydrated under
mechanical pressure. This characteristic of the hydrogels enables
them to creep, in the same way as the natural nucleus, under
compression and to withstand cyclic loading for long periods
without any significant degradation or loss of their elasticity.
This is because water in the hydrogel behaves like a cushion
whereby the polymeric network of a hydrogel with a high equilibrium
water content (EWC) is less susceptible to damage under mechanical
load.
Another advantage of hydrogels is their permeability to water and
water-soluble substances, such as nutrients, metabolites and the
like. It is know that body fluid diffusion, under cycle loading, is
the major source of nutrients to the natural disc. If the route of
this nutrient diffusion is blocked, e.g., by a water-impermeable
nucleus, further deterioration of the disc will ensure.
Hydrogels can be dehydrated and the resultant xerogels hydrated
again without changing the properties of the hydrogels. When a
hydrogel is dehydrated, its volume decreases, thereby facilitating
implantation of the prosthetic nucleus into the nuclear cavity in
the disc. The implanted prosthetic nucleus will then swell, in the
body, by absorption of body fluid up to its EWC. The EWC of the
hydrogel depends on the compressive load applied thereto. Thus, the
EWC of a specific hydrogel in an open container will differ from
the EWC of the same hydrogel in a closed vessel such as an
intervertebral disc. The EWC values, referred to below, are for
hydrogels subjected to compressive loads under the conditions found
in an intervertebral disc. The expansion factor of a dehydrated
hydrogel, in turn, is dependent on its EWC. Thus, it may vary from
1.19 for a hydrogel of 38% EWC to 1.73 for a hydrogel of 80% EWC.
For an 80% EWC hydrogel, the volume of the dehydrated prosthetic
nucleus is usually about 20% of that of the hydrated one. The
ability to be dehydrated and then return to its original shape upon
hydration, up to its EWC, makes it possible to implant the device
posterior-laterally during surgery, thereby reducing the complexity
and risk of intraspinal surgery as traditionally used. The danger
of perforation of the nerve, dural sac, arteries and other organs
is also reduced. In addition, the incision area on the annulus can
be reduced, thereby helping to heal the annulus and prevent the
reherniation of the disc. Hydrogels are also useful for drug
delivery into the disc due to their capability for controlled
release of drugs. Various therapeutic agents, such as growth
factors, long term analgesics and anti-inflammatory agents can
attach to the prosthetic nucleus and be released in a controllable
rate after implantation of the nucleus in the disc.
Furthermore, dimensional integrity can be maintained with hydrogels
having a water content of up to about 90%. This dimensional
integrity, if the nucleus is properly designed will aid in
distributing the vertebral load to a larger area on the annulus
ring and prevent the prosthetic nucleus from bulging and
herniating.
However, it is normally difficult to implant a fully hydrated
hydrogel prosthesis in the cavity, of a disc, through the small
window provided in the disc, for removing the herniated nucleus,
especially in a percutaneous surgery by virtue of their bulkiness
in a fully in a fully hydrated state. Therefore, such prosthesis
must be implanted, in the disc in relatively dehydrated states
which requires long periods to achieve their EWCs due to their low
surface areas. Other hydrogels, having high surface areas, do not
completely conform to the shape of the nuclear cavity. Other
polymers such as those disclosed in WO 97/268407 (PCT/US97 00457),
the teachings of which are incorporated herein by reference, can
also be used to fill the disc nucleus.
It is desirable to provide a hydrogel implant which is inherently
radiopaque, i.e., radiovisible so that surgeons could view the
placement of the implant in the cavity produced by the removal of a
spinal nucleus. It is advantageous if the radiovisible material
could be incorporated into the polymeric or hydrogel material
making up the prosthetic nucleus implant. It is desirable to have a
method of making the hydrogel or polymer radiopaque which would
allow dimensional changes in the hydrogel implant during processing
and after implantation without compromising the mechanical
integrity of the implant.
Various methods are used to implant a hydrogel or other polymeric
nucleus implant. Such a method is shown in U.S. Pat. No. 5,800,549,
the teachings of which are incorporated herein by reference.
SUMMARY OF THE INVENTION
It is an object of the invention to provide a novel hydrogel or
other polymeric implant for replacing the resected natural nucleus
of a spinal disc.
It is a further object of the invention to provide a novel hydrogel
or other polymeric prosthetic nucleus which has radiovisible
material contained therein.
It is yet a further object of the invention to provide a method for
incorporating radiovisible material within the hydrogel or polymer
either dispersed throughout the implant or in discreet locations
therein.
Such objects are achieved by the spinal implant for replacing the
natural nucleus of the disc made of a hydrogel having radiovisible
material located within the hydrogel. The material may be a metal
such as gold, tungsten, tantalum, platinum, titanium or
combinations thereof.
The material may be in powder form and may be distributed
throughout the hydrogel in a uniform manner or may be in powder
form and placed in the hydrogel or polymeric implant in discreet
layers or locations. The radiopaque powder preferably has a maximum
diameter of between 10 and 100 .mu.m and more preferably the powder
has a diameter of about 75 .mu.m.
Alternately, the metal may be in the form of foil, either in strip
form in the form of flakes scattered throughout the implant. If the
foil is in strip form, it should be relatively thin, i.e., in the
range of 1-100 .mu.m thick so that when used with a hydrogel, it
expands and contracts as the hydrogel is hydrated and dehydrated.
In the preferred embodiment, a thickness at the lower end of this
range is desirable, for example, 2 .mu.m thick.
In an additional embodiment, the implant may be in the form of a
metal wire or coil placed in the hydrogel implant during its
formation. Again, the wire is of such a diameter as to be able to
fold upon itself during hydration and dehydration of the
hydrogel.
If the polymer used is formed in situ then the metal that imparts
radiovisibility to the implant is dispersed throughout the polymer
prior to injection and curing.
If the polymer is processed in the melt then the metal is blended
into the polymer when it is above its melt temperature.
Methods of making the hydrogel, including the radiopaque material,
include dissolving the polymer powder to form a homogeneous
solution and then mixing the metal powder with metal flakes therein
while the solution is still a liquid and then freezing the solution
to form a solid. Usually the solution is poured into a mold to form
the hydrogel and then the mold is placed in the freezer.
Alternately, the hydrogel implant can be formed sequentially by
placing a homogeneous solution with radiopaque material in the mold
but filling only a portion of the mold, freezing the solution,
placing an additional layer of hydrogel, including the radiopaque
metal upon the first layer of solidified hydrogel and thereafter
freezing the second layer to form a solid. This process may be
repeated to form alternate layers in the prosthetic nucleus which
are either radialucent or radiopaque.
In yet another embodiment, the polymeric implant can be formed by
melting a polymer such as poly(acrylonitrile) and blending the
radiopacifying agent prior to molding the implant. Sequential
molding operations performed on one implant can result in the
radiopacifying agent being localized into discreet areas of the
implant. The polymeric implant can also be formed by injecting both
a crosslinkable material (e.g. monomer and/or prepolymer) and a
crosslinking agent, and then allowing the crosslinkable material to
cure in situ such as polyurethane. Radiovisibility can be imparted
to this implant by adding the radiopacifying agent to the
crosslinkable material, the crosslinking agent, or both.
If a coil or foil is used to produce the radiopaque portions of the
hydrogel, the foil or coil may be placed in the liquid hydrogel
prior to solidifying it by cooling.
BRIEF DESCRIPTION OF THE DRAWINGS
FIG. 1 is an elevational view of the vertebral disc absent its
nucleus with its associated vertebra;
FIG. 2 is an elevated view of the intervertebral disc and
associated vertebra of FIG. 1 from which the nucleus has been
removed;
FIG. 3 is an elevational view of the disc of FIG. 2 with the
polymeric nucleus of the present invention implanted therein
showing a radiopaque powder dispersed throughout the implant;
FIG. 4 is an elevational view of the disc of FIG. 2 with a
polymeric implant having a foil strip therein;
FIG. 5 is an elevational view of the disc space of FIG. 2 with a
polymeric implant having a metal coil therein; and
FIG. 6 is an elevational view of the disc of FIG. 2 having a
polymeric implant, including seven layers with three layers
including radiopaque metal powder dispersed therethrough.
DETAILED DESCRIPTION
Referring to FIGS. 1 through 6, in the preferred embodiment the
hydrogel prosthetic nucleus of the present invention, generally
denoted as 10, conforms when hydrated to its EWC, generally to the
shape of the natural nucleus. Alternately, the hydrogel can be
constrained in a polymer jacket. Such is taught in U.S. Pat. Nos.
5,674,295 and 6,132,465. The prosthetic nucleus is implanted within
cavity 11 the disc 12 of the vertebrae 14 and is surrounded by the
natural annulus 16. Vertebral end plates 20 and 22, as shown in
FIG. 1, cover the superior and inferior faces of nucleus 10
respectively. The implant is inserted through an opening 62 in
annulus 12.
Referring to FIG. 3, there is shown the prosthetic nucleus of the
present invention, including metal particles 30 dispersed uniformly
throughout. Uniformly is used in a relative sense not to indicate
that the particles are exactly spaced within the hydrogel.
Referring to FIG. 4, there is shown the prosthetic disc nucleus of
the present invention with a radiopaque foil strip 40 located
therein.
Referring to FIG. 5, there is shown a polymeric implant of the
present invention having a radiopaque or radiovisible coil 50
located therein.
Referring to FIG. 6, there is shown a prosthetic disc nucleus 10 of
the present invention having radiopaque layers of metal particles
or foil particles 60 located therein. While three layers are shown
in FIG. 6, one layer or two layers or more than three layers may be
utilized. As shown in FIG. 6, the three layers 60 having the
radiopaque powder or particles extend across the entire distance
between the inner walls of cavity 11 of annulus 16 of disc 12.
These layers are sandwiched between two layers having no radiopaque
particles. This structure results from the method of manufacturing
nucleus 10 as set forth in Example 2 below.
Hydrogels useful in the practice of the invention include lightly
cross-linked biocompatible homopolymers and copolymers of
hydrophilic monomers such as 2-hydroxylalkyl acrylates and
methacrylates, e.g., 2-hydroxyethyl methacrylate (HEMA); N-vinyl
monomers, for example, N-vinyl-2-pyyrolidone (N-VP); ethylenically
unsaturated acids, for example, methacrylic acid (MA) and
ethylenically unsaturated bases such as 2-(diethylamino) ethyl
methacrylate (DEAEMA). The copolymers may further include residues
from non-hydrophilic monomers such as alkyl methacrylates, for
example, methyl methacrylate (MMA), and the like. The cross-linked
polymers are formed, by known methods, in the presence of
cross-linking agents, such as ethyleneglycol dimethacrylate and
methylenebis (acrylamide), and initiators such as 2,2-azobis
(isobutyronitrile), benzoyl peroxide, and the like, and radiation
such as UV and .gamma.-ray.
Methods for the preparation of these polymers and copolymers is
well known to the art. The EWC of these hydrogels can vary, e.g.,
from about 38% for Polymacon.TM. (poly HEMA) to about 79% for
Lidofilcon.TM. B (a copolymer of N-VP and MMA) under ambient
conditions.
Another type of hydrogel, useful in the practice of the invention,
is illustrated by HYPAN.TM. and poly(vinyl alcohol) (PVA)
hydrogels. These hydrogels, unlike the aforementioned hydrogels,
are not cross-linked. Their insolubility in aqueous media is due to
their partially crystalline structures. HYPAN.TM. is a partially
hydrolyzed polyacrylonitrile. It has a multiblock copolymer (MBC)
structure comprising hard crystalline nitrile blocks, which provide
the hydrogel with good mechanical properties, and soft amorphous
hydrophilic blocks to provide the hydrogel with good water binding
capability. The methods of preparing HYPAN.TM. hydrogels of
different water contents and mechanical properties have been
disclosed in the U.S. Pat. Nos. 4,337,327, 4,370,451, 4,331,783,
4,369,294, 4,420,589, 4,379,874 and 4,631,188. The pre-nuclear
forms of this material, for use in this invention, can be prepared
by melt processing using solvents such as DMF and DMSO, as melting
aids or by solution processing.
Other types of polymers useful in the practice of the invention
include medical grade polyurethanes and materials formed by
crosslinking protein precursors. These materials may or may not
form hydrogels in their final form but are still useful as
materials to form prosthetic nucleus replacements. Such materials
are shown in U.S. Pat. Nos. 5,888,220, 6,189,048, 6,183,518 and in
Publication U.S. 20020049498 A1, the teachings of which are
incorporated herein by reference.
A preferred hydrogel for use in the practice of this invention is
highly hydrolyzed crystalline poly (vinyl alcohol) (PVA). The
amount of hydrolyzation may be between 95 and 100 percent depending
on the desired EWC which will be from about 60% to about 90%.
Generally, the final hydrogel water content increases with
decreasing hydrolyzation of the initial PVA which results in
decreased crystallinity.
Partially crystalline PVA hydrogels may be prepared, from
commercially available PVA powders, by any of the methods disclosed
in the U.S. Pat. No. 4,663,358, the teachings of which are
incorporated herein by reference. Typically, 10-15% PVA powder is
mixed with a solvent, such as water, dimethyl sulfoxide (DMSO),
ethylene glycol and mixtures thereof. A preferred solvent is 15%
water in DMSO. The mixture is then heated at a temperature of about
100 to about 120.degree. C., until a viscous solution is formed.
The solution is then poured or injected into a tubular metal, glass
or plastic mold and allowed to cool to below -10.degree. C.,
preferably about -20.degree. C.
The solution is maintained at the temperature for several hours,
preferably about 20 hours, during which time crystallization and,
therefore, gelation of the PVA occurs. The shaped gel is soaked
with several portions of water which are periodically replaced,
over a period of at least two days, until all the organic solvent
in the gel has been replaced by water. The hydrated gel can then be
partially or completely dehydrated for implantation. The hydrogels
thus prepared have EWC's between 60-90% and compressive strengths
of at least 1 MNm.sup.-2, preferably about 4 MNm.sup.-2, when
subject to the same constraints as the natural nucleus in an
intervertebral disc. In general, any polymer that can be used for
biomedical purposes can be used as long as the polymer exhibits the
desired stiffness characteristics.
Completion of the solvent exchange is determined by known methods.
For instance, when the solvent is DMSO its removal, from the gel,
is determined as follows:
50 .mu.L of a 0.01 N KMnO.sub.4 are added to 50 mL aliquots of the
water which has been separated from the gels. The presence of DMSO,
in the water, will be indicated by disappearance of the
characteristic pink color of the KMnO.sub.4. When the DMSO has been
completely removed, the pink color will not disappear. This method
has a detection limit of 0.3 ppm, for DMSO, when compared to a
blank and 0.3 ppm aqueous DMSO standard.
In general, any hydrogel that can be used for biomedical purposes
can be used as long as the hydrogel exhibits an EWC from about 30
to 90% and a compressive strength of at least about 1 MNm .sup.-2,
preferably 4 MNm.sup.-2, when subjected to the constraints of the
annulus and end plates of the disc. A rod or tube made from these
materials, in a dehydrated form, i.e., as xerogels, can be prepared
either by cast molding or lathe cutting. In cast molding, the
liquid monomer mixture, with initiator, is poured into a mold of
predetermined shape and size, and cured. If desired, the casting
mixture may include water, or another aqueous medium. Under those
circumstances the resultant rod or tube will be partially hydrated,
i.e., a hydrogel. In the case of lathe cutting, the xerogel can be
prepared, in a similar manner to the above, in the form of a block
or rod which is larger than needed to form the prosthetic nucleus.
The xerogel is then cut to the shape and size required for
implantation into the disc cavity. In both cases, the hydrogel
expansion factor, due to polymer swelling upon hydration, has to be
taken into account in designing the mold or in cutting the block,
rod or tube.
The exact size of the prosthetic nucleus, at its EWC, can be varied
for different individuals. A typical size of an adult nucleus is
about 2 cm in the semi-minor axis, about 4 cm in the semi-major
axis and about 1.2 cm in thickness.
Polymers curable within the body can also be used to replace the
natural nucleus and strengthen the annulus which is made of
cartilage. Natural cartilage is a non-vascular structure found in
various parts of the body. Articular cartilage tends to exist as a
finely granular matrix forming a thick incrustation on the surfaces
of joints. The natural elasticity of articular cartilage enables it
to break the force of concussions, while its smoothness affords
ease and freedom of movement. Preferred biomaterials, therefore,
are intended to mimic many of the physical-chemical characteristics
of natural cartilage. Biomaterials can be provided as one component
systems, or as two or more component systems that can be mixed
prior to or during delivery, or at the site of repair. Generally
such biomaterials are flowable in their uncured form, meaning they
are of sufficient viscosity to allow their delivery through a
cannula of on the order of about 2 mm to about 6 mm inner diameter,
and preferably of about 3 mm to about 5 mm inner diameter. Such
biomaterials are also curable, meaning that they can be cured or
otherwise modified, in situ, at the tissue site, in order to
undergo a phase or chemical change sufficient to retain a desired
position and configuration.
When cured, preferred materials can be homogenous (i.e., providing
the same chemical-physical parameters throughout), or they can be
heterogenous. An example of a heterogenous biomaterial for use as a
disc replacement is a biomaterial that mimics the natural disc by
providing a more rigid outer envelope (akin to the annulus) and a
more liquid interior core (akin to the nucleus). In an alternative
embodiment, biomaterials can be used that provide implants having
varying regions of varying or different physical-chemical
properties. With disc replacement, for instance, biomaterials can
be used to provide a more rigid, annulus-like outer region, and a
more fluid, nucleus-like core. Such di- or higher phasic cured
materials can be prepared by the use of a single biomaterial, e.g.,
one that undergoes varying states of cure, or by using a plurality
of biomaterials. Examples of suitable biomaterials includes, but
are not limited to, polyurethane polymers.
The in situ cured polymer may comprise a thermosetting polyurethane
polymer based on a suitable combination of isocyanates, long chain
polyols and short chain (low molecular weight) extenders and/or
crosslinkers. Suitable components are commercially available and
are each preferably used in the highest possible grade, e.g.,
reagent or preferably analytical grade or higher. Examples of
suitable isocyanates include 4,4'-diphenyl methane diisocyanate
("MDI"), and 4,2'-diphenylmethane diisocyanate ("TDI"). Examples of
suitable long chain polyols include tetrahydrofuran polymers such
as poly(tetramethylene oxide) ("PTMO"). Particularly preferred are
combinations of PTMO's having molecular weights of 250 and 1000, in
ratios of between about 1 to 1 and about 1 to 3 parts,
respectively. Examples of suitable extenders/crosslinkers include
1,4-butanediol and trimethylol propane, and blends thereof,
preferably used at a ratio of between about 1 to 1 and about 1 to 7
parts, respectively. Such performance can be evaluated using
procedures commonly accepted for the evaluation of natural tissue
and joints, as well as the elevation of biomaterials.
In particular, the in situ cured polymer forms, exhibit mechanical
properties that approximate those of the natural tissue that they
are intended to replace. For instance, for load bearing
applications, preferred cured composites exhibit a load bearing
strength of between about 50 and about 200 psi (pounds per square
inch), and preferably between about 100 and about 150 psi. Such
composites also exhibit a shear stress of between about 10 and 100
psi, and preferably between about 30 and 50 psi, as such units are
typically determined in the evaluation of natural tissue and
joints.
Biomaterials provided as a plurality of components, e.g., a
two-part polyurethane system, can be mixed with the radiopaque or
radiovisible metal powder at the time of use using suitable mixing
techniques, such as those commonly used for the delivery of
two-parts adhesive formulations. More preferably, the metal powder
can be added during melt processing of the polymer. A suitable
mixing device involves, for instance, a static mixer having a
hollow tube having a segmented, helical vein running through its
lumen. A two-part polyurethane system can be mixed by forcing the
respective components through a lumen, under pressure.
The hydrogels and polymers of the present invention have a much
higher structural integrity than the natural nucleus, i.e., they
are deformed with greater difficulty under a mechanical compressive
load (shaped gel vs. loose gel). That is because, unlike the loose
gel of the natural nucleus, the shaped gel has shape memory due to
the cross-linking or strong hydrogen bonding in the polymer matrix.
However, it would still have extensive lateral bulging under high
compressive load if there were no boundaries to constrain the
deformation. Because use of the prevent invention does not involve
removal of the disc annulus and/or end-plates, the lateral bulging
of the hydrogel nucleus will be restricted by the restoring forces
of the stretch fibers. Also, due to its superior structural
integrity, the hydrogel nucleus will not herniate or bulge through
the previously herniated areas or the incision which was made to
remove the degenerated nucleus.
If a hydrogel is used, since the natural nucleus is also primarily,
a hydrogel, the implanted prosthetic nucleus can easily restore all
the biomechanical functions of the nucleus which had been removed.
Unlike the prior art prosthetic discs, the hydrogel nucleus of the
present invention will restore the viscoelastic behavior of the
disc due to the water binding capability of the prosthetic
hydrogel.
The implantation of a prosthetic nucleus 10 can be performed in
conjunction with a discectomy or chemonuclealysis. Because the
properties of the prosthetic nucleus of the present invention are
similar to those of the nucleus material, the herniated nucleus can
be partially or totally replaced by the hydrogel prosthetic
nucleus. Due to the small size of the prosthetic it can be
implanted into the disc by means of a posterior lateral approach,
thereby significantly reducing the difficulty and the risk of the
operation.
The volume of a hydrogel nucleus of about 80% EWC will be reduced
by about 80% (to about 20% of its original volume) when dehydrated.
Consequently, the surgeon does not need to jack apart the vertebrae
adjacent to a damaged disc as required by, for example, the device
disclosed in U.S. Pat. No. 4,772,287. The height of the dehydrated
prosthetic nucleus, when inserted, is smaller than the disc space.
Furthermore, the rigidity of the dehydrated prosthetic nucleus will
help the surgeons to manipulate the prosthetic nucleus during the
operation. After implantation, the hydrogel nucleus of the present
invention swells in the body to a predetermined height which is
enough to maintain the space between the vertebral body. The
swelling process normally takes several hours to two days depending
on the size of the prosthetic nucleus and type of hydrogel.
The preferred method for making the radiopaque hydrogel or polymer
involves incorporating a metallic element into the structure of the
implant. The metallic element is in a form that will allow it move
with the polymeric structure as the implant changes dimensions
and/or geometry. This property is important because it minimizes
any internal stress amplification that could be caused by
incorporating the metallic component into the dynamic (i.e.
non-fusion) spine implant.
Three embodiments of the invention are described. The first
embodiment involves incorporating radiopaque materials such as a
metal powder, for example, gold or tungsten, into the polymer while
it is a liquid either due to the use of solvents, heat if the
polymer can be processed in the melt, or is in a pre-polymer form
prior to a curing step. The metal powder has a nominal diameter of
10-100 .mu.m, with a preferred maximum size of 75 .mu.m. The powder
is incorporated into the liquid polymer solution/melt preferably in
a concentration between 0.02 and 0.5 g per cc of polymer, with a
preferred concentration of 0.1 g/cc. The powder may be evenly
dispersed throughout the entire implant, or incorporated into the
implant in specific areas only. For example, by combining the use
of liquid-phase polymer that contains no metal powder in a mold
with liquid polymer that does contain metal powder it is possible
to form radiovisible areas of an implant in a variety of geometries
(e.g. lines, discs, spheres). An embodiment of the layered
polymeric implant has one or more planes in the implant comprised
of the radiovisible polymer. It is possible to incorporate 1 to 5
or more bands of powder-filled polymer in any plane across an
implant that also has polymer regions where no metal powder exits
using techniques that prevent the powder from migrating until the
liquid-phase polymer has formed a solid, as discussed below or
creating an even dispersion of metal powder throughout the entire
implant.
A second embodiment involves the use of a metal foil 0.001 to 0.1
mm thick for example gold, tantalum or tungsten foil. The foil can
be placed into the liquid polymer in the form of sheets or strips,
or it can be chopped into small pieces and incorporated into the
implant in the manner described for metal powders. Small pieces of
foil may provide an advantage for some polymer systems with a lower
viscosity in the liquid phase because each piece may have less mass
than a metal particle, which would result in less tendency for
migration through the liquid-phase polymer. In addition, the
geometry of a piece of foil with multiple irregular folds may also
have less tendency to migrate through liquid-phase polymer than,
for example, a smoother spherical particle. Examples include
suspending 1 to 5 or more strips of metal foil in any plane in the
implant, or incorporating 1 to 5 or more bands of polymer that
contain small pieces of chopped foil that may or may not be
"crinkled" as described above for powder, and creating an even
dispersion of pieces of chopped metal foil that may or may not be
"crinkled" throughout the implant.
A third embodiment involves the use of a metal wire or coil,
preferably 0.01 to 1.0 mm in length, for example, gold wire,
tungsten wire or platinum wire. The metal wire is suspended in the
liquid-phase polymer in such a manner that the wire will be
completely encapsulated by solid polymer at the end of the
manufacturing process. Alternatively, the wire can be formed into a
coil or other shape that may provide better radiographic
information about the implant and have less of an ability to
migrate through the solid polymer. Examples include suspending 1 to
5 or more pieces of wire in the liquid-phase polymer in such a way
as they will not be exposed to the surface of the implant, or
incorporating 1 to 5 or more bands of polymer that contain small
pieces of chopped wire or creating an even dispersion of pieces of
chopped metal wire throughout the implant using techniques as
described below.
EXAMPLE I
A PVA solution was formed by mixing 15 g of PVA powder (Kuraray 117
or equivalent), having a molecular weight about 78000 and about
99.7% hydrolysed (Cat. No. 15129, Polysciences Inc., Warrington,
Pa.), with 85 ml of a solvent comprising 15% water in DMSO. The
mixture was heated at about 110.degree. C. until a homogenous
viscous solution formed.
EXAMPLE II
0.1 gram of gold powder (maximum diameter 75 .mu.m) per cc of
liquid-phase of PVA solution of Example I were mixed. The two
ingredients were combined in the following manner to create a
metal-filled polymer solution. A plunger from a 5 cc first syringe
was removed and the first syringe was slowly filled half way with
PVA solution, 0.5 g of the gold powder was poured into the syringe.
The first syringe was then completely filled with PVA solution, and
the plunger replaced. A two-way luer connector was screwed onto the
tip of the first syringe and the connector was primed with PVA
solution from the syringe. The first syringe and a second syringe
of equal size were connected using the connector. The gold powder
solution from the first syringe was squeezed into the empty
syringe. This was repeated until the solution was uniformly mixed.
A third syringe was filled with PVA solution without any gold. 5 cc
of the PVA solution of the third syringe was injected into a
nucleus mold having a total volume of about 20 cc and a diameter of
about 1.5 cm. The mold was cooled down in freezer (4.degree. C.)
for about 20 minutes. 1 cc of gold powder-filled PVA solution was
injected on top of the cooled solution in the mold. 9 cc of the PVA
solution was slowly injected into the mold, on top of the gold
powder solution. The mold was again cooled down in the freezer for
about 15 minutes. About 5 cc of gold powder-filled solution was
again injected on top of the cooled solution in the mold to form a
second radiopaque layer. The mold was then completely filled with
the PVA solution from the third syringe.
EXAMPLE III
About 20 cc of the gold solution was prepared as described in
Example II using two 20 cc syringes and 2 grams of gold powder. The
metal filled polymer solution was slowly injected into a 20 cc
implant mold (for a #5 size implant), filling the mold completely.
The mold reservoir was capped so that the metal-filled PVA solution
will not leak out of the mold if the mold is inverted. The mold was
pressurized and placed in a Turbula Mixer which was placed into a
programmable freezer. The mixer used must be able to both rotate
and tip the mold during gelation of the metal filled PVA solution
to keep the powder in the solution uniformly distributed. The mixer
can rotate the mold about a central axis, tilt the mold back and
forth through a 90.degree. arc about an axis perpendicular to the
central axis. This kept the metal particles uniformly suspended in
the solution until it gelled.
EXAMPLE IV
The PVA solution prepared for the third syringe of Example II
without the gold powder was used in this Example. A strip of gold
metal foil was suspended in an empty mold using a thin monofilament
of nylon so that the strip is positioned in approximately the
center of the mold. The nylon filament was smooth and non-porous.
The mold was slowly filled with the PVA solution. The mold was then
placed in the freezer and after completing the process, the
monofilament was pulled out of the implant leaving the strip of
metal foil intact.
EXAMPLE V
The gold foil of Example IV was cut into small pieces using a food
processor or other convenient method to create metal flakes. 0.03 g
of metal flakes per cc of the liquid phase polymer of Example I. 20
cc of solution were placed in a mold and the solution was then
frozen as above to form the implant.
EXAMPLE VI
A gold coil was suspended in an empty 20 cc mold using a thin
monofilament so that the coil is positioned in approximately the
center of the mold. The filament was smooth and non-porous. A barb
fitting through a small loop at the top of the metal coil is a
convenient way to attach the filament and the coil. Slowly fill the
molds with the PVA solution of Example I. After completing the
gelation process by freezing, the monofilament was pulled out of
the implant leaving the gold coil intact.
After gelation, the implants were sterilized as described in my Ion
Treated Hydrogel copending application, U.S. Ser. No. 10/020,389
and then packaged.
The implants are formed into 10 different sizes by volume with a
size range of 1.1 to 5.2 cc for future implantation.
Although the invention herein has been described with reference to
particular embodiments, it is to be understood that these
embodiments are merely illustrative of the principles and
applications of the present invention. It is therefore to be
understood that numerous modifications may be made to the
illustrative embodiments and that other arrangements may be devised
without departing from the spirit and scope of the present
invention as defined by the appended claims.
* * * * *